More than a century ago, based on observations of the microvasculature, Thoma proposed that longitudinal tension controls vessel length. Since then, a number of studies have shown that blood vessels can remodel either physiologically or pathologically when exposed to altered mechanical environments. Arteries exposed to elevated flow (such as arteries upstream of arteriovenous fistulas (Holman, Surgery 26:889–917 (1949); Shenk et al., Surg. Gynecol. Obstet. 110:44–50 (1960)), collateral arteries carrying flow around an obstruction (Mulvihill, et al., N. Engl. J. Med. 104:1032 (1931)), and aortorenal bypass grafts (Stanley, et al., Surgery 74:931 (1973)) remodel (autoregulate) to increase their luminal diameter in response to increased flow as the result of vascular smooth muscle cell relaxation. In contrast, arteries experiencing reduced flow decrease luminal diameter.
Animal studies substantiate these clinical observations and suggest that vessels remodel so as to restore the wall shear stress to initial levels (Fung et al., J. Appl. Physiol. 70(6):2455–2470 (1991); Kamiya et al., Am. J. Physiol. 239(1):H14–21 (1980); Zarins et al., J. Vasc. Surg. 5(3):413–420 (1987)). Inflation of a tissue expander implanted within a rat hind limb over different periods of time ranging from 2 to 21 days increased the length of adjacent blood vessels 83±43%. Relatively slow expansion-induced lengthening (≦10% per day) did not diminish vessel patency, though more rapid expansion did substantially reduced patency (Stark, Plastic and Reconstructive Surgery, 30(4):570–578 (1986)).
However, the complex interdependence between components of the mechanical environment (e.g., pressure, shear, and strain) in vivo has hindered the identification of the specific mechanical stimuli responsible for remodeling. For example, by altering the viscosity of the perfusing medium, Melkumyants and coworkers have reported that by decoupling the effects associated with shear rate, ∂vz/∂r, (e.g., convection-enhanced transport and streaming potentials) and the wall shear stress, −μ∂vz/∂r, that acute autoregulation is a response to wall shear stress, not to flow rate per se (Melkumyants et al., Cardiovasc Res. 24(2):165–168 (1990)). Several widely used systems that expose cultured endothelial and smooth muscle cells to well-defined mechanical environments exist, but extrapolating results from cell culture models to vascular remodeling has proven to be problematic.
Traditional organ culture models employing excised vessels, such as human saphenous veins under static conditions, provide a well-defined chemical/biochemical environment and have been used to study the effects of pre-existing intimal hyperplasia, surgical preparations (Soyomo et al., Cardiovasc. Res. 27(11):1961–1967 (1993)), and specific biochemical factors, including bFGF (Soyomo et al., 1993) and ET-1 (Porter et al., J. Vasc. Surg. 28(4):695–701 (1998); Masood, et al. Brit. J. Surg. 84(4):499–503 (1997) on intimal hyperplasia. The inadequacy of these models is evidenced by the fact that vessels maintained under static conditions, even in the absence of known biochemical atherogenic stimuli, rapidly undergo pathological remodeling, including substantial intimal hyperplasia (Soyomo et al., 1993).
The atherogenic nature of traditional organ culture models appears to be at least partially due to the absence of physiologically relevant levels of mechanical forces. Porter and coworkers developed a crude, first-generation flow system by cutting an excised saphenous vein longitudinally and gluing the adventitial surface of the vein to the inside of a perfused Tygon tube (Porter et al., Cardiovasc. Res. 31(4):607–614 (1996)). The application of venous levels of pressure and flow-induced shear stress to excised human saphenous veins partially attenuated intimal hyperplasia associated with traditional organ culture, while arterial levels of pressure and shear stress completely abolished intimal hyperplasia (Porter et al., 1996). These results showed that, with a mechanically active environment, it was possible to maintain blood vessels in organ culture for weeks without pathological changes.
While the mechanical environments used in these studies were intended to mimic aspects of the arterial or venous circulation, they lacked many relevant mechanical features, including temporal variations, cyclic strains, as well as pressure drops across the vessel wall and the resulting transmural flow—each of which is a potentially important mechanical stimulus to blood vessels as summarized in reviews by, e.g., Gooch et al., Mechanical Forces: Their Effects on Cells and Tissues, Berlin, Springer, 182 (1997), and by Liu, Crit. Rev. Biomed. Eng. 27(1–2):75–148 (1999). Perfusion systems have been developed and used to provide a sophisticated mechanical environment by introducing pulsatile flow, cyclic flexure (Vorp et al., Ann. Biomed. Eng. 27(3):366–371 (1999)) and transmural pressure (Chesler et al., Am. J. Physiol. 277(5 Pt 2):H2002–2009 (1999)). These have been used to study the effects of the mechanical environment on gene expression (Vorp et al, 1999), endothelial cytoskeleton (Herman et al., J. Cell Biol. 105(1):291–302 (1987), lipid transport across the endothelium (Herman et al, 1987), and vasomotor responses (Labadie et al., Am. J. Physiol. 270(2 Pt 2):H760–768 (1996)).
Perfusion systems have also been used to investigate the effect of hydrodynamic forces on endothelial cells, with specific focus on the mechanisms by which endothelial cells perceive a mechanical stimulus and convert it to the initial biochemical response (i.e., mechanotransduction) (Gooch et al., Am. J. Physiol. 270(2 Pt 1):C546–51 (1996)), as well as the effect of biochemical pathways stimulated by fluid flow and mechanical forces on cellular proliferation (Gooch et al., J. Cell Physiol. 171(3):252–258 (1997); Gooch et al., Mechanical Forces: Their Effects on Cells and Tissues, 1997)) and susceptibility to viral infection. In addition, the effect of a hydrodynamic environment on the development of tissue-engineered cartilage has been investigated (Gooch, K., et al., “Mechanical Forces and Growth Factors,” in Frontiers in Tissue Engineering, (C. Patrick, A. Mikos, and L. McIntire, editors.) Pergamon, New York. p. 61–82 (1998)).
Vessel cultures have also been used to explore the molecular biology of vascular remodeling, both under static (Porter et al., 1998; Masood et al., 1997; Porter et al., Brit. J. Surg. 85(10):1373–1377 (1998); Porter et al., Eur. J. Vasc. Endovasc. Surg. 17(5):404–412 (1999)), and mechanically active environments (Chesler et al., 1999; Meng et al., 1999). One area in which the ex vivo vessel models have been particularly insightful is mechanical regulation of matrix metalloproteinases (MMPs), expression and activity (Vorp et al, 1999; Chesler et al., 1999; Meng et al., Exp. Mol. Pathol. 66(3):227–237 (1999); Mavromatis et al., Arterioscler. Thromb. Vasc. Biol. 20(8): 1889–1895 (2000)), and the role of MMPs in vascular remodeling (Porter et al., 1998; Porter et al., 1999; Loftus et al., Ann. N Y Acad. Sci. 878:547–50 (1999)).
Tenascin-C (TN-C) is large (>1000 kDa), disulfide-linked, hexameric extracellular matrix (ECM) glycoprotein that is prominently expressed during embryonic development, epithelial-mesenchymal interactions, wound healing, cancer, and notably, vascular disease (Mackie, Int. J. Biochem. Cell Bio. 29(10):1133–1137 (1997)), and is also subject to mechanical regulation. TN-C expression has been shown to be increased in rats and children suffering from pulmonary hypertension (Jones et al., J. Cell Sci. 112(Pt 4):435–445 (1999)), and under increased mechanical loading regimes, TN-C expression co-localizes with neointimal lesions expressing epidermal growth factor (EGF) and proliferating cell nuclear antigen (PCNA) (Jones et al., J. Cell Biol. 139(1):279–293 (1997); Jones et al., Circ. Res. 79(6):1131–1142 (1996)). The pro-proliferative role of TN-C is supported by in vitro studies that show TN-C acts as a survival factor for cultured smooth muscle cells (Cowan et al., Circ. Res. 84(10):1223–1233 (1999)). The majority of studies show that soluble, extracellular, and matrix factors regulate TN-C at the transcriptional level (Chiquet-Ehrismann et al., Bioessays 17(10):873–878 (1995)). In addition, targeted suppression of TN-C arrests progressive pulmonary hypertrophy in organ culture (Cowan et al., 1999). Taken together, these data strongly suggest that in the vessel wall the expression of TN-C is regulated by the mechanical environment, and the expression of this protein in turn is a key regulator of SMC proliferation and vascular remodeling.
Nevertheless, there is a sizable unmet demand for effective small-diameter vascular prostheses for use in coronary bypass surgery. Currently, the best replacements for occluded arteries are autologous arteries, which have a cumulative patency rate of 93% after 5 years (Lytle et al., J. Thorac. Cardiovasc. Surg. 89(2):248–258 (1985)). However, the number of expendable autologous arteries of appropriate dimensions for bypass grafts is severely limited, although there are numerous expendable arteries of smaller dimensions.
In animal studies where autologous tissue-engineered small-diameter vessels were evaluated in vivo, they performed much worse than an autologous vein would have (e.g., about half of the tissue-engineered vessels had decreased perfusion or loss of patency within 1 month (Niklason et al., Science 284(5413):489–493 (1999); Campbell et al., Circ. Res. 85(12):1173–1178 (1999)). Donor veins of appropriate dimensions are more readily available and are frequently used, but they have a substantially lower patency. Human saphenous vein grafts have a patency of ˜90% at early time points, and 81% after 1 year (Fitzgibbon et al., J. Am. Coll. Cardiol. 28(3):616–626 (1996)), but this has been reported to diminish to 45% after 5 years (Lytle et al., 1985).
Thus, the limited availability of suitable autologous arteries, coupled with the poor long-term patency of autologous veins, has led researchers to explore a number of approaches to create small-diameter vascular prostheses. These include using natural (Sandusky et al., J. Surg. Res. 58(4):415–420 (1995)) and synthetic polymeric materials (Smith et al., J. Med. Chem. 39(5):1148–1156 (1996); Uretzky et al., J. Thorac. Cardiovasc. Surg. 100(5):769–776 (1990)), pre-endothelializing existing types of polymer grafts in vitro (Stansby et al., Cardiovasc. Surg. 2(5):543–548 (1994); Stansby et al., Brit. J. Surg. 81(9):1286–1289 (1994)), and creating bioartificial or tissue-engineered blood vessels from cells and various support structures (Weinberg et al, Science 231:397–400 (1986); L'Heureux et al., J. Vasc. Surg. 17(3):499–509 (1993); L'Heureux et al., FASEB J. 12(1):47–56 (1998); Tranquillo et al., Biomaterials 17(3):349–357 (1996); Niklason et al., 1999); Shinoka et al., J. Thorac. Cardiovasc. Surg. 115(3):536–545 (1998)). While there are a number of different approaches to generating autologous tissue-engineered vessels in vitro, they all follow the same general paradigm: isolate specific cell types from blood vessels, expand these cells in vitro, and reassemble these cells into a tissue-engineered blood vessel—with the last step being the major challenge.
Many of these approaches yielded tissue-engineered arteries that grossly resemble native vessels, but in animal studies where tissue-engineered vessels generated in vitro were evaluated in vivo, their performance was inferior to that of autologous veins (Niklason et al., 1999; Campbell et al., 1999; Fitzgibbon et al., 1996). However, it was generally found that the performance of autologous blood vessels (whole vessels) was clearly superior to that of tissue-engineered blood vessels (prepared from only cells derived from the vessels).
There remains, however, a need in the art for a method or system by which a blood vessel can be harvested and used to direct the growth of an intact vessel ex vivo, wherein the newly formed vessel would be of sufficient size to permit the formation of a tissue-engineered vessel, which would be suitable for use as an arterial graft in vivo. Criteria for assessing the remodeled arteries relate both to the extent that the vessels grow ex vivo, and the degree that the remodeled arteries resemble healthy arteries of corresponding dimensions. Even modest increases in vessel dimensions would be potentially useful. Based on rough estimates using Poiseuille's law (i.e., vessels deform iso-volumetrically (Milnor, Hemodynamics, 2nd, 1989)), increasing the internal arterial diameter by 33% will increase the ability of that artery to carry blood by more than 200%. Poiseulle's law,       Q    =                  -                              πΔ            ⁢                                                  ⁢            P                                8            ⁢            μ            ⁢                                                  ⁢            L                              ·              r        4              ,relates the volumetric flow rate, Q, to the radius of a straight cylindrical tube of radius, r. Increasing the radius from 100% X to 133% X, increases flow from Y to 3.1 Y, a greater than 200% increase in flow.
In addition, in light of the foregoing and because blood vessels in vivo actively remodel (i.e., change size and/or composition) in response to chronic changes in the mechanical environment, the utilization of this ability of intact blood vessels to remodel supports the use of the system and methods of the present invention as a more effective and alternate approach to generating tissue-engineered blood vessels.